Positron emission tomography (PET), also known as PET imaging, is a type of nuclear medicine imaging that uses radioactive material, placed in a patient's body, to identify molecular activity and processes and, thus, assist in diagnosing disease, evaluating medical conditions, monitoring a patient's response to therapeutic interventions, etc. As shown in FIG. 1, a PET system 10 generally comprises an imaging device 12 that can detect radioactive emissions from the radioactive material (also known as radiopharmaceuticals or radiotracers) in the internal body area P under examination, a data processor 14 that analyzes the detected emissions information, and an image processor 16 (which in some configurations may be part of the data processor 14) that converts the processed data into image data or images of the area under examination via mathematical image reconstruction software. A user interface 18 (which typically includes an associated display such a touch screen display or non-touch screen display, keyboard and/or mouse) accompanies the processors 14, 16 and controls the operation of the system 10 and the various components. Although not shown in detail, the various components are operably connected to one another via appropriate control circuitry which is manipulated via the user interface 18.
The imaging scanner 10 comprises a number of detector assemblies 30 of a scintillation crystal 32 optically coupled with a photosensor 34, for example, a photomultiplier tube (PMT) or avalanche photodiode (APD), that are arranged to form a series of concentric rings 40 (although other shapes, like hexagons or partial rings, may be formed). This is shown in FIGS. 2a and 2b. The detectors assemblies 30 (or scintillation detectors) are configured generally alike and may take on various coupling configurations, such as, one crystal 32 to one photosensor 34, a plurality of crystals 32 to one photosensor 34, or one crystal 32 to a plurality of photosensors 34. Both the crystals 32 and the photosensors 34 may be packed in modular structures and together in detector blocks 36. The scintillation detectors 30 are oriented in a respective detector ring 40 so the crystals 32 face the common central opening to form the inner diameter and the photosensors 34 are located behind the crystals 32 and form the outer diameter. The scintillation detectors 30 are further arranged in a respective ring 40 so that each has a diametrically opposed scintillation detector 30 across the central opening. The rings 40 are dimensioned so that the length of a patient's body P may be accommodated by the common central opening, in its axial direction (as shown in FIG. 1). Each detector ring 40 thus lies in a transverse plane of the patient's body P.
In operation, after an appropriate radiotracer is placed into a patient's body P and becomes concentrated in tissues of interest, the patient is placed in the central opening of the scanner 12. The radiotracer undergoes positron emission decay and each emitted positron travels in the tissue for a short distance until it interacts with an electron. The encounter annihilates both electron and positron, producing a pair of annihilation (gamma) photons γ moving in approximately opposite directions. The two photons γ travel to respective scintillation detectors 30 that are diametrically opposed within the detector ring 40. Each photon γ first enters and travels through the scintillation crystal 32 (the scintillator) which converts the high-energy photons into visible light (i.e., “optical light” or “optical photons”). The signal response of the scintillator 32 is typically a prompt intensity increase at the time of excitation followed by a decay with time. This is the scintillation process. The photosensor(s) 34 of the respective scintillation detector 30 detect the burst of light incident upon its coupling face(s) and, in turn, converts the light to an electrical signal.
All photons γ that interact with the detectors 30 are “registered” by the scanner 12 which forwards the electrical signals of registered events to the data processor 14. The data processor 14 analyzes the signals and determines if two registered events selected are a so-called coincidence event (i.e., a simultaneous or coincident detection of a photon pair). For this analysis, the data processor 14 utilizes only true coincidence events which occur when both photons γ of an annihilation event are detected by two opposing detectors 30 in coincidence within the resolving time of the scanner 12 (also known as the coincidence time-window), typically on the order of nanoseconds, and neither photon γ has undergone any form of interaction prior to detection. The source of a temporal photon pair then may be localized (via appropriate analysis) along the straight line that joins the two detectors 30 of the coincidence event (i.e., a line of coincidence or a line of response (LOR)).
Thousands of coincidence events are produced by each scan. Consequently, for each detector ring 40 (or transverse plane), a fan-beam of LORs over many angles by all the allowed coincident detector 30 pairs is produced. The large majority of coincidences are formed by diametrically-opposed detectors 30 in one ring 40 (“direct planes”) and a smaller percentage of those coincidences are formed between adjacent planes (“cross planes”). The smaller/faster the resolving time of the detectors 30, the more precise the localization can be (e.g., a segment of a chord) and the better the signal-to-noise ratio (SNR) of the image, requiring fewer events to achieve the same image quality.
The data processor 14 forwards all coincidence data to the image processor 16 where final image data is produced via mathematical image reconstruction algorithms and software. Briefly, a map of the sources of each temporal photon pair may be constructed using the thousands of coincidence events and solving sets of simultaneous equations for the total activities along the LORs. The resulting map shows the tissues in which the radiotracer has become concentrated, and can be interpreted by an appropriate health professional.
In any imaging system, such as a PET system 10, knowing the depth-of-interaction (DOI), i.e. the depth within the detector 30 the signal of interest interacted, improves the accuracy in determining the direction from which the signal originated. Knowing the DOI also permits correction of event timing information, i.e. when the signal of interest interacted within the detector 30. For a PET scanner 12, knowing the DOI improves the image resolution and the system timing. As noted above, an improvement in system timing yields improved SNR in the tomographic images, particularly in time-of-flight (TOF) systems. Knowing the DOI can also improve energy resolution if there are depth-dependent losses of signal intensity which also aids in improving final image quality.
There is much interest in DOI determination for scintillation detectors 30 and there are many proposed methods in the medical literature. These proposed methods include, for example:                using layers of dissimilar materials, such as scintillators of differing decay time, light yield efficiency, or wavelength to identify the layer of interaction;        using weighted light-sharing across boundaries orthogonal to the DOI sought, such that the projected 2D image (flood image, position profile) is distorted by the shared light as a function of DOI;        using photosensors at opposite ends of a scintillator and measuring the relative light intensity;        measuring the spread of the light, and mapping that to the DOI; and        placing multiple layers of thin detectors, either intrinsic or scintillation with an associated photosensor.        
However, these proposed methods suffer from complexity and cost, and introduce degradations in other PET system performance criteria, such as sensitivity, energy resolution and, often, timing. Hence, they are only seen in academic prototypes and commercial animal imaging systems.
In scintillation detectors 30, the optical light is isotropically generated and there are usually a limited number of coupling or sensor surfaces of the photosensor(s) 34 by which the light is detected. In order to get a reasonable estimate of the energy deposited within the scintillator 32, reflectors are placed on all non-sensing surfaces of the photosensor(s) 34. The fact that the light undergoes multiple and varied reflections, dependent on the location of the energy deposition, means that there will be an associated shape to the response signal (typically, energy distribution versus photon count rate) from the photosensor(s) 34 which is superimposed on the scintillation light's intrinsic decay time. Advantageously, this shape may be exploited to yield an estimate of the DOI.